Gradient coils

Abstract: A magnetic resonance imaging machine includes a toroidal vacuum dewer (24) which contains a superconducting magnet (10). A radio frequency coil (32) is mounted within a cylindrical bore (26) of the vacuum dewer. A cylindrical, dielectric former (46) supports an RF shield (34), a z-gradient coil (50), an x-gradient coil (52), and a y-gradient coil (54). The x and y-gradient coils are each composed of four like spiral coil constructions. A metallic layer is cut with cut lines (64) to define a generally spiral coil winding pattern. In a high current density region (68) in which the coil windings are narrower than a preselected width, the cut lines (76) are thinner. In lower current density regions (70), the cut lines (78) are thicker. In lower current density regions, two cut lines are defined between adjacent coil windings such that the coil windings are limited to a maximum width. Conductive islands (80) defined in the lower current density areas which are large enough to support detrimental eddy currents are preferably removed or at least modified such that their ability to support detrimental eddy currents is curtailed.

Description: This invention relates to gradient coils. It finds particular application in conjunction with sheet or "fingerprint" type gradient coil designs for magnetic resonance imaging apparatus and will be described with particular reference thereto. However, it will be appreciated that the present invention will also find application in conjunction with the generation of magnetic fields and gradients for other applications. Heretofore, magnetic resonance imagers have included a superconducting magnet which generated a temporally constant primary magnetic field through a central bore. A series of annular gradient magnetic field coils for generating x, y, and z-gradient magnetic field gradients were mounted to an interior of the bore. An annular radio frequency coil was commonly positioned in the interior of the gradient coils for transmitting radio frequency signals into and receiving radio frequency magnetic resonance signals from a subject in the bore. Current pulses were applied to the gradient and radio frequency coils to generate a series of RF and gradient field pulses of conventional magnetic resonance imaging sequences. Various coil constructions have been used for generating the x, y, and z magnetic field gradients. One type of gradient coil includes a flexible, dielectric backing layer to which a sheet of copper or other conductive foil material has been laminated. A coil pattern was defined by cutting, milling, or etching the conductive sheet in a generally spiral-like pattern. The conductive sheet remaining between the spiral cut provided a generally spiral or fingerprint-like current path. See, for example, U.S. Patent No. 5,177,442 to Roemer or U.S. Patent No. 4,840,700 to Edelstein. The conductive pattern represents an approximation of a continuous current density vector J. Mathematically, the current density J=curl S, where S is a stream function representing contours of constant integrated current density. A number of turns N is selected and the coil is patterned into N+1 contours, each of constant S, which differ in magnitude by an amount ΔS=S max /N. Hypothetical contours offset from these by ΔS/2 represent a desired pattern for placing filamentary wires to approximate J. The contour lines generated by this method determine the machining pattern or cut lines. The cut lines were formed by removing a constant width of the conductor material producing an electrical discontinuity or gap which defines the turn pattern. Typically, the conductive sheet is copper and has a thickness between 1 and 2 mm. The cut lines are typically 2 mm. wide, i.e. about equal to the sheet thickness. Because the current density function J varies across the sheet, this process of removing a constant width cut line to define the coil pattern results in conductors of varying width. The current can spread in the wide areas and becomes concentrated in the narrow conductor areas such that the selected current density J is better met. One of the difficulties which arises is that in regions of high current density, only narrow segments of conductor remain between the cut lines, e.g. a width of about 4 mm. This reduction in conductor width from a larger nominal width, e.g. 6 mm, is disadvantageous. The power loss, I 2 R, increases markedly with the increase in current density in the narrow regions. Moreover, the heat generation rate ρJ 2 (where ρ is electrical resistivity), also increases markedly in the high current density regions. Reducing the conductor width from 6 mm. to 4 mm. increases the power loss by a factor of about 1.5 and increases the heat generation rate by a factor of about 2.25. The localized coil heating is a significant disadvantage in the prior art. Another disadvantage of the prior art is that the wider portions of the conductor support eddy currents. The gradient coil commonly includes a layered assembly of x, y, and z-coils. All three are mounted in very close proximity to one another and are driven in a pulsed manner. When a coil is driven with a current pulse, the resultant magnetic field induces eddy currents in neighbouring conductors wherever possible. These eddy currents reduce the driving magnetic field and have an associated decay time. Because the neighbouring coils are also fingerprint coils, the available current paths for eddy currents have a non-constant spatial distribution. Thus, the eddy current patterns are spatially dissimilar from the driving field. Both the extended duration of the time decay and the spatial dissimilarity are disadvantageous in magnetic resonance applications. In addition, the electric field generated by the differences in voltages between adjacent coils results in capacitive coupling. Charging this capacitance also has an associated time constant. It is usually desirable to minimize the time constants and currents associated with this process as well. Wire wound gradient coils also have disadvantages. The wire width is limited by the minimum contour separation distance and the wire height/thickness is limited by radial build constraints. The smallest cross-section is used to wind the entire spiral coil. This results in a coil of higher resistance and higher total heat generation. The present invention provides a new and improved gradient coil configuration. According to a first aspect of the present invention there is provided a magnetic resonance imaging apparatus comprising: a generally toroidal magnet assembly for generating a generally uniform magnetic field through an examination region along the z-axis of an xyz co-ordinate system of the imaging apparatus, the magnet assembly including a generally cylindrical member which defines a bore which extends along the z-axis around the examination region; a gradient coil assembly for causing magnetic field gradients within the bore along the x, y and z axes, a radio frequency coil assembly, disposed within the bore for transmitting radio frequency pulses into the examination region for inducing and manipulating magnetic resonance of selected dipoles within the examination region; a sequence control means for controlling the gradient coil assembly and the radio frequency coil assembly for generating magnetic resonance imaging gradient and RF pulse sequences; an image reconstruction means for reconstructing image representations from the magnetic resonance signals emanating from the examination region, the gradient coil assembly including: a gradient coil including four electrically interconnected first coil constructions, two on each side of a central xy plane through the examination region, each first coil construction being disposed adjacent said central plane and extending circumferentially about half way around a cylinder, a second gradient coil including four second coil constructions, two on each side of said central plane, each second coil construction being disposed adjacent the central plane and extending circumferentially about half way around the cylinder, the second gradient coil being rotated 90° around the cylinder from the first gradient coil, each of the first and second coil constructions including an electrically conductive sheet which is divided by cut lines into a generally spiral electrically conductive winding pattern, the winding pattern having at least one high current density region in which windings of the winding pattern have a width below a preselected width, windings in lower current density regions of the winding pattern being at least the preselected width, characterised in that the cut lines are narrower in the high current density region than in the lower current density regions. According to a second aspect of the present invention there is provided a gradient coil assembly including a plurality of electrically interconnected coil constructions, each of the coil constructions including an electrically conductive layer which is divided by cut lines into a generally spiral electrically conductive winding pattern, the winding pattern having at least one high current density region in which windings of the winding pattern are narrower than a preselected minimum width, the windings in other regions of the winding pattern being at least the preselected minimum width, characterised in that the cut lines are narrower in the high current density region than in the other regions. According to a third aspect of the present invention there is provided a method of forming a gradient coil construction from a conductive layer laminated to a dielectric substrate, the method comprising: selecting a generally spiral current winding pattern centroid; comparing a distance between adjacent centroids with a first preselected width, in regions in which the distance is narrower than the first preselected width, defining a cut line through the conductive layer with a first cut line width and in regions in which the distance exceeds the first preselected width, defining a cut line of a second cut line width between adjacent centroids, the first cut line width being smaller than the second cut line width. An advantage of an apparatus and method embodying the present invention is that heating is reduced in regions of high current density. Another advantage of an apparatus and method embodying the present invention is that it reduces power loss in the gradient coils. The invention will now be described, by way of example, with reference to the accompanying drawings, in which:- FIGURE 1 is a diagrammatic illustration of a magnetic resonance imaging system in accordance with the present invention; FIGURE 2 is a top plan view of a quarter of an x or y-gradient coil laid out flat; FIGURE 3 is an enlargement of a section of the coil construction of FIGURE 2; and FIGURE 4 is a cross-sectional view of an alternative gradient coil to that of FIGURE 2. With reference to FIGURE 1, a plurality of superconducting magnetic field coils 10 generate a generally temporally constant main magnetic field along a longitudinal or z-axis of a central bore 12. The x and y-axes are defined orthogonal to the z-axis and each other, e.g. horizontally and vertically. Coils of the superconducting magnet are supported by a former 14 and received in a toroidal helium vessel or can 16. The helium vessel is filled with liquid helium to maintain the superconducting magnet at a temperature at which it is superconducting. A main magnetic field shield coil assembly 18 generates a magnetic field which opposes the field generated by the superconducting magnet 10 in regions surrounding the superconducting magnet. To reduce helium boil-off, the toroidal helium vessel 16 is surrounded by a first cold shield 20 which is maintained at about 20 K or less. A second cold shield assembly 22 is chilled to about 70 K or less. The first and second cold shields are preferably cooled by mechanical refrigeration (not shown). Preferably, the cold shield 22 includes a slotted tube cylindrical cold shield portion which is mounted on a cylindrical dielectric former. A toroidal vacuum vessel 24 encases the cold shields to define a vacuum reservoir therearound. The vacuum vessel 24 includes a cylindrical, non-ferrous member 26 that defines the bore 12 and extends parallel to the z-axis. The main magnetic field is generated through the cylindrical member 26. Preferably, layers of aluminized (aluminum coated) mylar superinsulation (not shown) are arranged between the vacuum vessel and the cold shields. An RF coil assembly 30 is mounted in the bore. The RF coil assembly includes a primary RF coil 32 and an RF shield or secondary coil 34. The primary coil 32, preferably a quadrature or birdcage-type coil, has an inner diameter which defines a patient receiving bore, preferably of at least 55 cm. A gradient coil assembly 40 includes a primary whole body gradient coil assembly 42 and a secondary coil assembly 44. The primary gradient coil assembly 42 includes a cylindrical dielectric former 46. Preferably, the former has a sufficient thickness to maintain its structural integrity under the forces associated with generating the magnetic field gradients, e.g. 5-10 mm of glass reinforced plastic, preferably about 7.5 mm. The former 46 has a series of grooves defined or cut in its outer surface to receive the windings of a z-gradient coil 50. The RF shield 34, preferably lapped strips of copper foil or fine copper mesh, is applied to the outer surface of the former 46 and through the z-coil receiving grooves. The RF shield mesh may be deformed into the grooves by tightly wound z-gradient wire windings, may be fitted into the grooves, may have solder connected segments extending through the grooves, or the like. The grooves have a depth which generally corresponds to the thickness of the z-gradient coils such that after the z-gradient coils are inserted, a substantially circularly cylindrical surface is presented for receiving x and y-gradient coils 52, 54. Various alternate embodiments are also contemplated. First, the primary z-gradient coil 50, may be either a distributed or bunched coil. Second, the coil may be constructed with hollow conductor for circulating a cooling fluid. This would enhance cooling efficiency so as to allow higher power levels or gradient strengths. Finally, the z-gradient coil need not be next to the RF screen but can be somewhat removed. With continuing reference to FIGURE 1 and particular reference to FIGURE 2, the x and y-gradient coils 52, 54 each include four windings of the pattern shown in FIGURE 2. Two of the constructions are mounted with edges 56 generally adjacent each other and on a vertical centre plane 58 of the bore. The two constructions wrap about half way around the dielectric former 46. An analogous pair of constructions are mounted on the opposite side of the dielectric former with their edges 56 again adjacent the central vertical plane 58. The constructions wrap around the former such that edges 60 of the oppositely disposed constructions are adjacent each other. The x and y-gradient coils are substantially equivalent to each other, but one is rotated 90° around the central or z-axis relative to the other. The four coil constructions of one of the gradient coils, e.g. the x-gradient coil, are mounted to the dielectric former 46 around the RF shield 34 and the z-gradient coil 50 and the assembly potted in epoxy. The four coil constructions of the other gradient coil, e.g. the y-gradient coil, are mounted around the potted x-gradient coil and also potted in epoxy or other dielectric material. Preferably, the constructions of the outer mounted gradient coil are extended between edges 60 by an appropriate percentage to compensate for the larger diameter about which it is mounted. With continuing reference to FIGURE 2 and further reference to FIGURE 3, a desired current density J is calculated for each coil construction and a centroid 62 of the current density J is determined. The calculated spiral centroid marks the winding pattern which an ideal filament coil would follow. Cut lines 64 are defined between adjacent centroids. A width of the conductive strip defined as the distance 66 between two adjacent centroids is determined and compared with a preselected minimum width, e.g. 7 mm. and a preselected maximum width, e.g. 12 mm. Where the conductive strips would have a width 66 of less than the preselected minimum, a region 68 of high current density is defined. Where the width 66 would be above the preselected maximum, a region 70 of low current density is defined. With continuing reference to FIGURES 2 and 3 and further reference to FIGURE 4, in the preferred embodiment, the coil construction includes a dielectric layer 72 on which a 2 mm. thick film 74 of copper is layered. In the high current density regions 68, narrow cut lines 76 are defined, preferably by a high pressure water jet cutting operation or other numerically controlled machining operator. In the preferred embodiment, the narrow cut lines 76 are about 1 mm. wide or less. In the regions 68 of higher current density, the width 66 of the copper foil windings is 6 mm. or less. At this narrow width, 1 mm. cut lines leave a significantly higher percentage of conductive material than would 2 mm. cut lines. In regions in which the width 66 is between 7 and 12 mm., wide cut lines 78 are cut with widths of about 2 mm. Although these regions could also be cut with cut lines of 1 mm., 1 mm. cutting operations can be significantly more expensive than 2 mm. cutting operations. In the regions 70 of low current density, the cut line spacing is limited to a preselected width, e.g. 12 mm., centre to centre. This limits the width of the conductive portion of the coil to 12 mm. less the 2 mm. cut line thickness, i.e. 10 mm. Preferably, the cut lines are positioned such that the width of the conductive coil segment is as close to 12 mm. as possible throughout the coil. When the width between cut lines is limited in the regions of low current density, islands 80 are formed. Preferably, the islands 80, or at least the larger islands, are removed. Very small islands, which are smaller than the width of the conductors, do not support significant eddy currents and can be retained. Optionally, rather than removing the larger islands, the larger islands may be cut into smaller pieces to prevent their maintaining large eddy currents. Also, it should be noted that one can vary the minimum and maximum prescribed widths depending upon the specific application. Various, additional techniques may be adopted for limiting eddy currents. First, as illustrated in FIGURE 4, a second foil layer 82 may be mounted to the opposite side of the dielectric layer. With two conductive layers, each layer may be half as thick. Hence, any eddy currents induced in the conductors will have a shorter time constant. Moreover, the winding pattern of the two conductive layers need not be the same. Rather, the current density J may be calculated with twice as many, i.e. 2N, loops with alternate loops being defined on each face. Alternate ones of the N loops can be formed on opposing faces allowing each loop to be twice as wide in the high current density region. As yet another alternative, the second conductive layer 82 may be defined only in the region 68 of high current density. The cut lines are directly over top of each other, such that analogous conductive strips are formed on both sides of the high current density region. Electrical connectors, such as copper pins 84, are drilled through the dielectric layer to interconnect the upper and lower conductive strips. The top and bottom conductive strips need not be the same thickness. Rather, the thickness of one or both of the layers can be adjusted such that the combined current carrying capacity of the upper and lower strips matches the current carrying capacity of strips in other regions of the coil construction. The use of thinner conductors are advantageous in that they facilitate defining the cut lines by photoetching techniques. As yet another alternative, the thickness of the conductive layer may be adjusted to be thinner in the areas of less current density and thicker in the areas of higher current density. To the extent that conductor portions would be so wide that they might support eddy currents, additional cut lines can be used to divide these portions into a plurality of parallel conductors connected electrically in parallel. With reference again to FIGURE 1, a plurality of plastic I-beams or other mechanical structures 90 mount the gradient former 46 to the interior surface of the cylindrical member 26 of the superconducting magnet. The I-beams or other mechanical supports 90 define a sufficient gap for receiving main magnet field shims (not shown), typically a multiplicity of ferrous elements, for shimming the uniformity of the main magnetic field. Preferably, an annulus of about 33 mm is provided between the primary gradient coil former 46 and the cylindrical member 26. This provides passages for air cooling, displaces the primary and shield gradient coils a maximal amount, and provides a compartment within which to insert magnetic field shims for shimming the uniformity of the main magnetic field. The secondary gradient coil 44 includes the cylindrical member 26 which preferably has grooves for receiving a z-gradient shield coil 92. X and y-gradient shield or secondary coils 94, 96 are mounted on the outer circular periphery of the former 26 and potted in an epoxy compound. The primary and secondary gradient coils are connected in series, or independently driven, to create accurate gradient magnetic fields in the bore 12 and cancelling magnetic fields outside the bore, particularly in the vacuum vessel. A sequence control means 100 controls a gradient coil control means 102 and a transmitter 104. The gradient coil control 102 is connected with a series of current pulse generators 106 which, in turn, are connected with the primary gradient coils 50, 52, 54 and the secondary gradient coils 92, 94, 96. The transmitter 104, preferably a digital transmitter, is connected with the primary radio frequency coil for generating pulses of radio frequency signals for exciting and manipulating magnetic resonance in selected dipoles of a portion of a subject within the bore. A radio frequency receiver 108, preferably a digital receiver, is connected with the radio frequency primary coil or surface coils (not shown) for demodulating magnetic resonance signals emanating from the examined portion of the subject. An image reconstruction means 110, such as an inverse two-dimensional Fourier transform reconstruction means, reconstructs the received magnetic resonance signals into an electronic image representation that is stored in an image memory 112. A video processor 114 converts electronic images stored in the memory 112 into appropriate format for display on a video monitor 116.

Claims: A magnetic resonance imaging apparatus comprising: a generally toroidal magnet assembly (10,14,16,18,20,22,24,26) for generating a generally uniform magnetic field through an examination region (12) along the z-axis of an xyz coordinate system of the imaging apparatus, the magnet assembly including a generally cylindrical member (26) which defines a bore which extends along the z-axis around the examination region (12); a gradient coil assembly (40) for causing magnetic field gradients within the bore along the x, y and z axes, a radio frequency coil assembly (30) disposed within the bore for transmitting radio frequency pulses into the examination region (12) for inducing and manipulating magnetic resonance of selected dipoles within the examination region (12); a sequence control means (100) for controlling the gradient coil assembly (40) and the radio frequency coil assembly (30) for generating magnetic resonance imaging gradient and RF pulse sequences; an image reconstruction means (110) for reconstructing image representations from the magnetic resonance signals emanating from the examination region (12), the gradient coil assembly (40) including: a gradient coil (52) including four electrically interconnected first coil constructions, two on each side of a central xy plane (58) through the examination region (12), each first coil construction being disposed adjacent said central plane (58) and extending circumferentially about half way around a cylinder (46), a second gradient coil (54) including four second coil constructions, two on each side of said central plane (58), each second coil construction being disposed adjacent the central plane (58) and extending circumferentially about half way around the cylinder (46), the second gradient coil (54) being 15 rotated 90° around the cylinder (46) from the first gradient coil (52), each of the first and second coil constructions including an electrically conductive sheet which is divided by cut lines (64) into a generally spiral electrically conductive winding pattern, the winding pattern having at least one high current density region (68) in which windings of the winding pattern have a width below a preselected width, windings in lower current density regions (70) of the 20 winding pattern being at least the preselected width, characterised in that the cut lines (64) are narrower in the high current density region (68) than in the lower current density regions (70). A magnetic resonance imaging apparatus according to claim 1 wherein the conductive sheet is a foil of constant thickness, the cut line width in the high current density region (68) being half the foil thickness and the cut line width in the other regions of the coil assembly being at least the foil thickness. A magnetic resonance imaging apparatus according to claim 2 wherein the foil thickness is at least 2 mm. A magnetic resonance imaging apparatus according to claim 1 wherein the coil construction has regions of lower current density in which the windings are wider than in the high current density region, the width of the windings having a maximum width in the lower current density region. A magnetic resonance imaging apparatus according to claim 4 wherein the windings have a substantially constant width in the regions of lower current density. A gradient coil assembly including a plurality of electrically interconnected coil constructions, each of the coil constructions including an electrically conductive layer which is divided by cut lines (64) into a generally spiral electrically conductive winding pattern, the winding pattern having at least one high current density region (68) in which windings of the winding pattern are narrower than a preselected minimum width, the windings in other regions of the winding pattern being at least the preselected minimum width, characterised in that the cut lines are narrower in the high current density region than in the other regions. A method of forming a gradient coil construction from a conductive layer laminated to a dielectric substrate, the method comprising: selecting a generally spiral current winding pattern centroid (62); comparing a distance between adjacent centroids with a first preselected width, in regions in which the dimension distance is narrower than the first preselected width, defining a cut line through the conductive layer with a first cut line width and in regions in which the dimension exceeds the first preselected width, defining a cut line of a second cut line width between adjacent centroids, the first cut line width being smaller than the second cut line width. A method according to claim 7 further including comparing inter-centroid dimension with a second preselected width, when the inter-centroid dimension exceeds the second preselected width, defining two cut lines between adjacent centroids, each cut line being displaced from a corresponding adjacent centroid by substantially one half the second preselected width, whereby islands of conductive layer material are defined between adjacent windings. A method according to claim 8 wherein the first and second preselected widths are the same. A method according to claim 8 further including removing the islands to prevent eddy currents from being supported therein. A method according to claim 8 further including modifying the islands to reduce the island's ability to support eddy currents.

Inventors: DEMEESTER GORDON D, MORICH MICHAEL A, PATRICK JOHN L, DEMEESTER, GORDON D., MORICH, MICHAEL A, PATRICK, JOHN L.
Applicants: MARCONI MEDICAL SYS INC, MARCONI MEDICAL SYSTEMS, INC.
Application year: 1994
Main classification: G01R33
IPC codes: G01R33